We developed a new reflection-mode optical-resolution photoacoustic microscopy (OR-PAM) based on the cooperation of a reflective objective and an ultrasonic transducer. The reflective objective is used to produce nearly diffraction-limited optical focusing, and the excited ultrasound waves are then directly detected by an ultrasonic transducer that was placed in the central cone of the objective. This new design avoids the coupling between optical focusing and ultrasound transmission in the reflection mode. Moreover, the proposed system is able to provide lateral resolution of 1.2 μm at 580 nm, penetration depth of 0.9 mm in biological tissues, and a work distance of 6.0 mm. We present in vivo imaging of the microvasculature in mouse ears and in vitro imaging of red blood cells (RBCs), which demonstrate the capability of the system to study microcirculation.
© 2013 Optical Society of America
Photoacoustic (PA) imaging is a promising hybrid-modality imaging technique, which can map the optical energy deposition in biological tissues . Due to the rich optical absorption contrast in biological tissues, this technique can provide structural, functional and molecular imaging of tissues [2–4]. Photoacoustic microscopy (PAM) is a newly developed imaging method facilitated by optical and ultrasonic focusing [5–8]. It can be classified into two categories: OR-PAM and acoustic-resolution PAM (AR-PAM) [1, 9]; they determined the lateral resolution by optical focusing and acoustic focusing, respectively. Since the optical focusing is more effective, OR-PAM provides higher resolution than AR-PAM to image microvasculature at capillary level with a shallow imaging depth [8, 10–17]. Moreover, OR-PAM can perform functional imaging of microvasculature, such as imaging of total hemoglobin concentration, hemoglobin oxygen saturation, metabolic rate of oxygen, blood flow velocity, and other functional parameters [18–23].
There are two detection configurations in OR-PAM system: transmission mode and reflection mode. In transmission mode, optical excitation and ultrasonic detection are located on the opposite sides of the sample, and the objective lens with high numerical aperture (N.A.) can be conveniently adopted to obtain high lateral resolution in sub-wavelength level. In reflection mode, optical excitation and ultrasonic detection are located on the same side of the sample. This configuration makes the system much more flexible in imaging of different imaging objects than transmission mode, especially with some thick samples. In the reflective mode, the optical-acoustic combiner is always used to deliver laser to samples and redirect the excited ultrasound waves to the ultrasonic transducer. However, this design is prone to deteriorate both the focusing of laser and the transmission efficiency of ultrasound waves, due to the coupling of laser and ultrasound [8, 10, 24]. To address this issue, several improved designs were proposed in the most recent works. Yao et al. utilized a ring-shaped focused ultrasonic transducer and a water-immersion objective lens to develop a high-resolution reflection-mode UV-PAM . But the size of the hole on the ultrasound transducer for the transmission of light and the focal length of the ultrasound transducer restrict the improvement of the resolution. Zhang et al. employed a special-fabricated parabolic ultrasonic mirror with a central hole to separate the delivering of laser and ultrasound wave, achieving a high-resolution reflection-mode OR-PAM . But the introduction of the ultrasonic mirror limited the working distance of the system. While in some in-vivo imaging, such as brain imaging, the long working distance with high resolution is favorable. These systems may not be able to fulfill the requirement. Therefore, further efforts are still required to obtain the OR-PAM with both high resolution and long working distance.
In this paper, we developed a new reflection-mode OR-PAM system with high resolution in visible optical spectral range. In the system, a reflective objective that has a long working distance and a large NA was used to achieve nearly diffraction-limited optical focusing. Moreover, a full size ultrasonic transducer using polyvinylidene fluoride (PVDF) polymer was placed in the central cone of the objective to directly detect the excited ultrasound waves. The lateral resolution, axial resolution, imaging depth, and the working distance of the system are evaluated in our study. Furthermore, in vitro imaging of RBCs and in vivo imaging of the microvasculature in mouse ear are presented.
Our OR-PAM system shown in Fig. 1 employs a reflective objective for optical focusing and a PVDF ultrasonic transducer for direct detection of the ultrasound waves excited by laser pulses. The system utilizes a wavelength tunable laser system, including a diode-pumped solid-state Nd:YLF laser (IS8II-E, Edgewave) and a dye laser (Credo, Sirah). Firstly, the laser beam from the dye laser is reshaped by an iris. Then, the reshaped beam is focused by a condenser lens. A pinhole (P50C, Thorlabs) with a diameter of 50 μm is well located near the focus of the beam as a spatial filter. The filtered beam is attenuated by a neutral density filter and collimated by another convex lens to fit for a single-mode fiber coupler (F-91-C1, Newport, with a 4 × objective). The beam from the single-mode fiber (P1-460A-FC-2, Thorlabs) is collimated by an aspheric lens (AL2550-A, Thorlabs) before being reflected by a mirror to fill the back aperture of the reflective objective (NT68-188, Edmund) for optical focusing. A beam sampler is inserted into the output beam of the laser to introduce some energy to a fast photodiode for triggering the data acquisition. The reflective objective has a long working distance of 23.2 mm and a nominal N.A. of 0.5. A customized meniscus lens with a hole in the center is used for correction of the optical aberration caused by air-water interface refraction. The task of the correction lens is making the light traveling along its former direction. The N.A. of the objective can be calculated with the formula N.A. = n × sin α, where n is the refractive index of the media between objective and the sample, and α is the angle aperture of the objective. When the objective is used in the air without correction lens, the refractive index of the media is 1. When the objective is used with the correction lens, the refractive index of the media changed to 1.33, and the α stayed still. Therefore, the N.A. of our detector is 1.33 times higher than it before, which is 0.665. The excited PA waves are received by a PVDF based ultrasonic transducer (PT50-3, Japan Probe; center frequency, 45 MHz; focal length, 10 mm), which is located in the central cone of the objective and immerged into water through the central hole of the correction lens. The objective and the ultrasonic transducer are aligned coaxially to achieve the overlap of the focus of the objective and the transducer for maximum sensitivity.
During the imaging, the imaging head shown in the dashed box in Fig. 1 is moved by a two-dimensional (2-D) high-precision linear stage (ANT95-50, Aerotech) to perform the raster scan. The position signals from the controller of the stage are used to trigger the laser output during the scanning, which is conducted at each assigned position. The excited PA signals at each position is amplified by two low-noise amplifiers (ZFL-500, Mini-Circuits), and then acquired by a 12-bit high-speed data acquisition card (ATS9350, Alazar Tech) at a sampling rate of 500 MHz. An imaging window is opened at the bottom of the water tank and is sealed with a polyethylene membrane for optical and ultrasonic transmission. Ultrasound gel is used for ultrasound coupling between the membrane and the tissue before imaging. Because the PA signals are depth resolved, the information of depth locations of the targets in tissue can be obtained by multiplying the delay time and the sound velocity of the tissue. Therefore, a 2-D raster scanning of the imaging head can form a three-dimensional (3-D) volumetric image of the tissue. We can also obtain a 2-D image by the maximum amplitude projection (MAP) along the axial direction.
The BALB/c mice weighting 20 to 25 g, obtained from the Animal Biosafety Level III Laboratory (Wuhan, China) and housed under specific pathogen-free (SPF) conditions, were used in the experiments. All experimental procedures are approved by Institutional Animal Ethics Committee of Huazhong University of Science and Technology. The mice were intraperitoneally anesthetized with a mixture of a-chloralose and urethane, and then the hairs on the ear were removed by using commercial human hair-removing lotion before imaging. Body temperature of the mice was kept constant at 37°C by utilizing a feedback controlled heating pad during imaging. The in vitro RBCs were also derived from the BALB/c mice. We imaged the single RBCs in blood smear made by the diluted blood containing anticoagulant.
3.1 Resolution evaluation of the system
Lateral resolution of the OR-PAM system is determined by the nearly diffraction-limited optical focal diameter of the objective. A sharp edge in an Air Force resolution test target (USAF-1951, Edmund) in water was imaged to quantify the lateral resolution at the light wavelength of 580 nm. After scanning across the edge with a step of 0.2 μm, the PA amplitude values were fitted by a sigmoidal-shaped function as the fitted edge spread function (ESF) of the system . The fitted ESF is shown in Fig. 2(a), and the line spread function (LSF) shown in Fig. 2(b) is calculated from the ESF. The full width at half-maximum (FWHM) of the LSF was used to estimate the system lateral resolution. In this way, the system lateral resolution was estimated to be 1.2 μm.
According to the aforementioned parameters of the objective, the diffraction-limited theoretical lateral resolution is about 0.45 μm. The experimental measured value is worse than the result of theoretical calculation. This is mainly due to the imperfect correction of optics refraction at the interface of air and water of the correcting lens. Furthermore, the reflective objective block the central part of the incidence beam, which makes the optical energy in zero-order diffraction bright spot is decreased. This may also slightly degrade the resolution of our system.
The axial resolution of our system was measured using 0.5 μm carbon nanoparticles. A PA axial spread profile of a typical 0.5 μm carbon nanoparticle is shown in Fig. 2(c). Because the size of the particle is much smaller than axial resolution, the axial spread profile can be regarded as the axial point spread function of the imaging system. The axial resolution of the system can be estimated by measuring the FWHM of the envelope of this axial profile . The envelope was achieved by adopting the absolute value of Hilbert transform of the PA signal. As shown in Fig. 2(c), the axial resolution was quantified as 30 μm in this way, agreeing with the value calculated from the measured bandwidth of the ultrasonic transducer (45 MHz in receiving-only mode) and the average sound velocity (1.5 mm/μs) in biology tissues.
3.2 Measurement of the penetration depth of the system
We measured the maximum penetration depth of our OR-PAM system by laying a black tape obliquely into chicken’s breast. A human hair closely onto the tissue surface was used as a reference. The imaging was taken with a laser wavelength of 580 nm, and the pulse energy out of the objective is about 60 nJ. The focus was located about 0.4 mm below the reference surface to perform scanning along the hair. As shown in Fig. 3(a), the black tape is clearly identified down to 0.9 mm below the surface with a signal-to-noise ratio (SNR) higher than 6 dB. Another test was also performed. Several human hairs were inserted into the chicken’s breast at different depths, and a few hairs were placed on the top surface of the tissue. The cross-section images of the hairs were obtained by scanning across the hairs. As shown in Fig. 3(b), a hair down to 0.92 mm beneath the surface is still visible with good SNR (> 6 dB). Therefore, the penetration depth of the OR-PAM system was estimated to be better than ~0.9 mm in biological tissues. Simultaneously, the working distance of the system was determined to be 6.0 mm by measuring the distance between focal plane and optical-acoustic probe.
3.3 In vivo imaging of microvasculature in a mouse ear
The wavelength used in the in vivo imaging is 584 nm, which is an isosbestic absorption wavelength of oxy- and deoxy-hemoglobin. Therefore, the image reflects the relative total hemoglobin concentration in blood vessels. An imaging area with the size of 2 mm × 3.4 mm at the end of a mouse ear was selected. As shown in Fig. 4(a), there are no large blood vessels visible with naked eyes in this area. Then the imaging area was scanned using our OR-PAM with a step size of 2 μm. Considering the repetition rate of the laser pulse is 1 kHz, this imaging can be accomplished in about 29 minutes. The 3-D volumetric rendering of the imaged microvasculature in the imaging area is shown in Fig. 4(b), where the morphology of the microvasculature is clearly displayed. The 2-D MAP image of the same area is also shown in Fig. 4(c). As shown in the image, a pair of concomitant venule and arteriole labeled with an arrow and their further complicated and detailed branches are clearly identified. Figure 4(d) shows a close-up MAP image in the area indicated by a box in (c) with a projection depth of 50 μm under the surface of the ear, which displays much smaller micro blood vessels. Many capillaries with diameter of 6-8 μm are visible in the image, and a capillary bed including thick capillaries indicated by a circle is also imaged with good contrast. After the imaging, there is no visible optical damage to the ear. Moreover, an area with the size of 0.5 mm × 1.2 mm close to the root of the mouse ear was also imaged with our OR-PAM using a step size of 1 μm × 1 μm. And the data acquisition time of this area is about 10 min. The MAP image of the microvasculature of this area is shown in Fig. 5(a). It can be seen that the blood vessels with different diameters in the imaged volume can be visualized with good contrast. To highlight the smaller blood vessels in superficial layer of the ear, a MAP image of micro blood vessels in superficial layer with a thickness of 90 μm of the ear is also shown in Fig. 5(b). Many capillaries with diameters of 5-8 μm are clearly displayed in the image, and the partial capillary network is also mapped with good quality. The obtained SNR of the image is 30 dB. Figure 5(c) shows the 3-D view of three MAP images from three parallel layers with a thickness of 60 μm beneath the skin surface from the same volumetric data as the image 5(a), and the layer spacing is 60 μm. Figure 5(d) shows the close-up of the single capillaries enclosed by the box in 5(b), where these discrete points may be RBCs. Although RBCs are the dominate contrast in the capillaries in the wavelength we used, we failed to conclude that they are RBCs, since we cannot image them with a donut structure, limited by the resolution of our system. In this experiment, the frame rate of B-scan is less than 1fps. Considering the moving speed of these suspicious RBCs in capillary is normally much higher than 1μm/s (the step size of imaging is 1μm), only a slice of the fast moving RBC can be acquired in this case. This makes some ‘horizontal pattern’ appearing in the Fig. 5(d).
3.4 In vitro imaging of RBCs
In order to validate the capacity of the system to image individual RBCs, we conducted PA imaging of the in vitro RBC sample. As shown in Fig. 6(a), individual RBCs can be clearly resolved using our OR-PAM system, the size of most cells is 5-6 μm. An image from transmission optical microscope (0.3 NA) of a mouse blood sample with the same place is shown in Fig. 6(b) for comparison. It is worth noting that, the distribution of RBCs in Fig. 6(a) and (b) seems quite different. This is mainly because the depth of the focus of optical microscope is about 7μm, while the axial resolution of the PAM is about 30μm. When we acquire the PAM image in Fig. 6(a), it is the projection of the maximum intensity along the depth direction. Therefore, the difference of imaging depth and depth of focus of PAM and optical microscope makes the images looks quite different.
We developed a reflection-mode high-resolution OR-PAM system in visible optical spectral range based on a reflective objective. The lateral and axial resolution was measured to be 1.2 μm and 30 μm, respectively. The imaging depth in chicken’s breast is measured to be better than 0.9 mm. We also validate our system in both in vivo imaging of mouse ear and in vitro imaging of RBCs.
Owning to the special structure of reflective objective, it is feasible for us to fit a normal ultrasonic transducer into the cone of objective. This design effectively separates the focusing of the laser and the transmission of the ultrasound, resulting in both nearly diffraction-limited optical focusing and direct detection of ultrasound waves. Furthermore, our system preserves the working distance as 6 mm. This will make our system more practicable in many biological applications. Besides, since the reflective objective can be used in a quite wide spectrum, this design will facilitate the functional imaging, especially when we want to use visible light to measure the hemodynamic parameters and detect the molecular information with IR laser at the same time.
It is worth noting that we cannot find a commercialized reflective objective for water immersion usage, and we have to make a correction lens to couple the transmission of the laser from air to water. The bad performance of this lens degrades the resolution of our system seriously. Furthermore, since the design of the reflective objective block some low-frequency part of the incident laser beam, the intensity of the beam center at the focal spot will be lower than that in the normal Gauss beam, and the side lobe of the focal spot will be enhanced. Therefore the weakening of the main lobe and the enhancement of the side lobe of the focal spot will both further deteriorate the resolution of the system. And our future work is to carefully design a water immersion reflective objective to improve the performance of this design.
The authors would like to thank Guoqiang Xu, Zhen Wang, Prof. Zhihong Zhang for preparation of samples. This work was supported by National Major Scientific Research Program of China (Grant No. 2011CB910401), Science Fund for Creative Research Group of China (Grant No. 61121004), and National Natural Science Foundation of China (Grant No. 81201067).
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